EP2508842B1 - Verfahren und System zur optischen Kohärenztomographie - Google Patents

Verfahren und System zur optischen Kohärenztomographie Download PDF

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Publication number
EP2508842B1
EP2508842B1 EP11002894.1A EP11002894A EP2508842B1 EP 2508842 B1 EP2508842 B1 EP 2508842B1 EP 11002894 A EP11002894 A EP 11002894A EP 2508842 B1 EP2508842 B1 EP 2508842B1
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Prior art keywords
dimensional
images
depth
coherence tomography
image
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EP11002894.1A
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German (de)
English (en)
French (fr)
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EP2508842A1 (de
Inventor
Rainer Nebosis
Roland Reuter
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Agfa HealthCare NV
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Agfa HealthCare NV
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Priority to EP11002894.1A priority Critical patent/EP2508842B1/de
Priority to ES11002894.1T priority patent/ES2497190T3/es
Priority to US14/009,379 priority patent/US9551565B2/en
Priority to PCT/EP2012/001436 priority patent/WO2012136339A1/de
Priority to CN201280017175.2A priority patent/CN103443578B/zh
Publication of EP2508842A1 publication Critical patent/EP2508842A1/de
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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01B—MEASURING LENGTH, THICKNESS OR SIMILAR LINEAR DIMENSIONS; MEASURING ANGLES; MEASURING AREAS; MEASURING IRREGULARITIES OF SURFACES OR CONTOURS
    • G01B9/00—Measuring instruments characterised by the use of optical techniques
    • G01B9/02—Interferometers
    • G01B9/0209—Low-coherence interferometers
    • G01B9/02091—Tomographic interferometers, e.g. based on optical coherence
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/0059—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence
    • A61B5/0062—Arrangements for scanning
    • A61B5/0066—Optical coherence imaging
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/0059—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence
    • A61B5/0073—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence by tomography, i.e. reconstruction of 3D images from 2D projections
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/44—Detecting, measuring or recording for evaluating the integumentary system, e.g. skin, hair or nails
    • A61B5/441—Skin evaluation, e.g. for skin disorder diagnosis
    • A61B5/444—Evaluating skin marks, e.g. mole, nevi, tumour, scar
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01B—MEASURING LENGTH, THICKNESS OR SIMILAR LINEAR DIMENSIONS; MEASURING ANGLES; MEASURING AREAS; MEASURING IRREGULARITIES OF SURFACES OR CONTOURS
    • G01B9/00—Measuring instruments characterised by the use of optical techniques
    • G01B9/02—Interferometers
    • G01B9/02055—Reduction or prevention of errors; Testing; Calibration
    • G01B9/02062—Active error reduction, i.e. varying with time
    • G01B9/02063—Active error reduction, i.e. varying with time by particular alignment of focus position, e.g. dynamic focussing in optical coherence tomography
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01B—MEASURING LENGTH, THICKNESS OR SIMILAR LINEAR DIMENSIONS; MEASURING ANGLES; MEASURING AREAS; MEASURING IRREGULARITIES OF SURFACES OR CONTOURS
    • G01B9/00—Measuring instruments characterised by the use of optical techniques
    • G01B9/02—Interferometers
    • G01B9/02083—Interferometers characterised by particular signal processing and presentation
    • G01B9/02085—Combining two or more images of different regions
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01B—MEASURING LENGTH, THICKNESS OR SIMILAR LINEAR DIMENSIONS; MEASURING ANGLES; MEASURING AREAS; MEASURING IRREGULARITIES OF SURFACES OR CONTOURS
    • G01B9/00—Measuring instruments characterised by the use of optical techniques
    • G01B9/02—Interferometers
    • G01B9/02083—Interferometers characterised by particular signal processing and presentation
    • G01B9/02089—Displaying the signal, e.g. for user interaction
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
    • G01N21/17—Systems in which incident light is modified in accordance with the properties of the material investigated
    • G01N21/47—Scattering, i.e. diffuse reflection
    • G01N21/4795—Scattering, i.e. diffuse reflection spatially resolved investigating of object in scattering medium
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B2560/00—Constructional details of operational features of apparatus; Accessories for medical measuring apparatus
    • A61B2560/04—Constructional details of apparatus
    • A61B2560/0437—Trolley or cart-type apparatus
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B2562/00—Details of sensors; Constructional details of sensor housings or probes; Accessories for sensors
    • A61B2562/02—Details of sensors specially adapted for in-vivo measurements
    • A61B2562/0233—Special features of optical sensors or probes classified in A61B5/00

Definitions

  • the present invention relates to a method and a corresponding system for optical coherence tomography.
  • OCT optical coherence tomography
  • biological tissue Due to its light-scattering properties, biological tissue is particularly suitable for the diagnostic examination by OCT. Since OCT operates with relatively low light intensities and the wavelengths of the light used are mostly in the near infrared range (750 nm to 1350 nm), in contrast to ionizing X-ray diagnostics for biological tissue, it does not represent a radiation load. It is thus of particular importance for medicine and roughly comparable to ultrasound diagnostics, where light is used instead of sound in OCT. The transit times of the light reflected at different boundary layers in the sample are detected by means of an interferometer.
  • OCT optical coherence tomography
  • At least two two-dimensional output images of an object are recorded in mutually spaced, mutually parallel planes of the object by means of an optical coherence tomography device, the output images each comprising a multiplicity of output image values, in particular intensity values. Further, an interpolation of the output image values of the at least two two-dimensional output becomes images are performed in three-dimensional space, whereby interpolation values are obtained which form a two-dimensional final image.
  • the inventive system for optical coherence tomography comprises an optical coherence tomography device for recording at least two two-dimensional output images of an object in mutually spaced, mutually parallel planes of the object, wherein the output images each comprise a plurality of output image values, in particular intensity values, and comprises processing means for interpolation the output image values of the at least two two-dimensional output images in three-dimensional space, whereby interpolation values are obtained which form a two-dimensional final image.
  • the output image values of the at least two two-dimensional output images lie in a regular grid in three-dimensional space, and the invention is characterized in that adjacent output image values in the grid have the same spacings in all three spatial directions.
  • the invention is based on the idea of using the optical coherence tomography device initially to record two output slices in two parallel planes of the object and then to interpolate the output image values of the two output slices in three-dimensional space, which is spanned by the two spaced-apart slices, so that a two-dimensional End section image from interpolation values, ie interpolated output image values.
  • output image values are interpolated not only of an output slice but of both output slice images.
  • An interpolation value is derived from at least one output image value of a first output slice image and at least one output image value of a second output slice image.
  • an interpolation value is derived from eight output image values, with four output image values from the first output slice image and four output image values from the second output slice image.
  • cavities or other structures with a size of more than approximately 10 ⁇ m are more apparent than is the case with the respective initial sectional images. But even smaller structures can be identified more reliably and more quickly and possibly examined. Particularly in the field of dermatology, it is thereby possible to recognize and examine certain diagnostically relevant or interesting structures in the skin with particularly high reliability.
  • the output image values of the at least two two-dimensional output images lie in a regular grid in three-dimensional space, wherein adjacent output image values have the same distances in all three spatial directions, in particular between approximately 2 ⁇ m and 4 ⁇ m.
  • the interpolation according to the invention of the output image values of the two output images can thereby be realized in a simple manner.
  • the output image values of the at least two two-dimensional output images are interpolated by a trilinear interpolation and / or a tricubic interpolation.
  • Trilinear or tri-cubic interpolation is a method of interpolating in a three-dimensional regular grid, i. a grid with the same lattice constant in all three spatial directions, wherein an interpolation value lying in the center of the respective lattice cells of the lattice is determined from eight output image values which lie at the eight corners of the lattice cell by linear or cubic interpolation.
  • the output image values located at the eight corners of the grid cell represent the nearest neighbors of the respective interpolation value, which is why the trilinear or tricubic interpolation can also be understood as a nearest-neighbor interpolation.
  • the trilinear or tricubic interpolation can also be understood as a nearest-neighbor interpolation.
  • a quadratic interpolation can also be carried out.
  • the output image values of the at least two two-dimensional output images are then interpolated by tri-quadratic interpolation.
  • the at least two two-dimensional output images of the object are real-time images taken at a rate of at least one image per second, preferably at least five images per second.
  • the interpolation according to the invention represents a simple and very fast method which can be applied to real-time recorded OCT images without a time delay.
  • output images recorded in real time can also be reproduced on a corresponding display device in real time, ie with a corresponding image repetition rate, in the form of end images interpolated according to the invention with correspondingly better diagnostic information.
  • the at least two two-dimensional output images of the object are taken in spaced-apart planes of the object in a first operating mode, in which reflected by the object or backscattered light from only a partial area, in particular of two adjacent rows of a spatially resolving detector the optical coherence tomography device is detected while the optical distance of a reflector is changed to a beam splitter of the optical coherence tomography device by an optical path which is substantially, in particular at least 100 times greater than the average wavelength of coupled into the optical coherence tomography light.
  • This mode of operation allows the acquisition of output images in the form of slices, at high speed and consequently in real time, i.e., in real time. at a rate of at least one image per second, easily and with high reliability.
  • the increased diagnostic validity of the real-time images acquired and interpolated in this way allows an even more reliable examination of the object.
  • the at least two two-dimensional output images of the object will be received in spaced-apart planes of the object in a second mode of operation, during a change in the optical distance of a reflector to a beam splitter of the optical coherence tomography device reflected from the object Light is detected by detector elements of a detector several times, in particular at most five times, wherein the change in the optical distance of the reflector to the beam splitter is at most forty times the average wavelength of coupled into the optical coherence tomography light.
  • output images in the form of two-dimensional tomograms, so-called en-face images, with high repetition rate, in particular in real time are recorded.
  • the two planes of the object preferably extend at different depths above or below a mean depth in the object.
  • the two planes extend at an equal distance above or below the mean depth in the object.
  • the distance of the two planes relative to each other preferably corresponds to the same distance of the output image values in all three spatial directions.
  • the mean depth or the different depths of the planes extending above or below the mean depth in the object are adjusted by the distance of the reflector from the beam splitter.
  • the optical distance of the reflector to the beam splitter of the optical coherence tomography device is changed by an optical path which is substantially, in particular at least 100 times greater than the average wavelength of the light coupled into the optical coherence tomography device.
  • the number of original output image values of the at least two two-dimensional output images in at least one dimension, in particular in the direction of the depth of the object is reduced by at least two, preferably more than ten, original output image values be summarized to an output image value.
  • the original output image values are samples obtained by successively sampling an interference pattern obtained from different depths within a depth range of the object.
  • those original output image values which are obtained from an area, in particular the depth area, of the object are summarized, whose extent in the at least one dimension of the resolution, in particular the axial resolution or depth resolution corresponding to the optical coherence tomography device.
  • these measures take into account that the sampling of the interference signal must be high enough so as not to violate the so-called sampling theorem.
  • this also takes into account the fact that the distance between two samples of the interference signal is generally much smaller than the physical resolution of the imaging optics of the optical coherence tomography device. This means that several, preferably more than ten, consecutive samples contain approximately the same physical information and can therefore be combined into one value without significant loss of information.
  • the output image value corresponds, for example, to an average value or the median from the original output image values.
  • the depth of field of the imaging optics of the optical coherence tomography device is greater than the spatial separation of the output image values from one another, which is preferably identical in all three spatial directions.
  • FIG. 1 shows a schematic representation of an example of an optical coherence tomography device, hereinafter also referred to as OCT device, with an interferometer 10, which comprises a beam splitter 11, a lighting arm 12, a reference arm 13, a sample arm 14 and a detector arm 15. Further, a radiation source 21 is provided for generating light, which is filtered by an optical filter 22 and is focused by an optics composed of lenses 23 and 24 to an input region 25 of a light guide 26. The radiation source 21 together with the optical filter 22 forms a device, which is also referred to as light source 20.
  • OCT device optical coherence tomography device
  • the light coupled into the light guide 26 is coupled into the illumination arm 12 of the interferometer 10 through an optical system 28 located in its output region 27. From there, the light first passes to the beam splitter 11, by which this on the one hand forwarded in the reference arm 13 and reflected by a movable reference mirror 16 located at its end and on the other hand illuminated after passing through the Probenarms 14 a surface 2 of a sample 1.
  • a plurality of in a, preferably flat, surface arranged detector elements comprises and consequently a spatially resolved detection of the light reflected from the sample 30 or a corresponding interference pattern due to its superposition with the light reflected at the reference mirror 16.
  • the detector 30 used is preferably a CMOS camera whose detector elements (so-called pixels) are sensitive in the infrared spectral range, in particular in a spectral range between approximately 1250 nm and 1350 nm.
  • the CMOS camera preferably has 512 ⁇ 640 detector elements.
  • the light guide 26 is preferably a so-called multimode fiber whose numerical aperture and core diameter allow not only one fiber mode to be formed at a certain wavelength of the light coupled into the fiber, but many different fiber modes can be excited.
  • the diameter of the multimode fiber used is between about 1 mm and 3 mm, in particular about 1.5 mm.
  • the size of the illuminated surface 2 on the sample 1 corresponds approximately to the size of the illuminated surface 17 on the reference mirror 16 and on the one hand by the optics located at the input region of the light guide 26, which in the example shown, the lenses 23 and 24 comprises, and on the other hand the intended in the exit region of the optical fiber 26 optics 28 determined.
  • the resulting interference pattern is detected by the detector 30, wherein a corresponding interference signal is generated.
  • the sampling rate of the detector 30 for sampling the interference signal must be chosen so that the time variation of the interference pattern can be detected with sufficient accuracy. This generally requires high sampling rates when high speeds are to be achieved for a depth scan.
  • a depth scan is preferably realized in the system described in that the optical distance of the reference mirror 16 is changed by the beam splitter 11 during the detection of the reflected light from the sample 1 with the detector 30 at a speed v by an optical path which is substantially larger as the mean wavelength of the light coupled into the interferometer 10.
  • the light reflected in at least 100 different depths of the sample 1 is preferably detected by the detector 30.
  • the optical path is changed periodically with an amplitude which is substantially greater than the average wavelength of the light coupled into the interferometer 10.
  • the change of the optical distance of the reference mirror 16 about the optical path or with the amplitude is preferably at least 100 times, in particular at least 1000 times, greater than the average wavelength of the light coupled into the interferometer 10. Due to the large paths in this distance variation, this movement of the reference mirror 16 is also referred to as a macroscopic movement.
  • the maximum possible scanning speed in the direction of the depth of the sample 1 is dependent on the maximum possible scanning rate of the detector 30.
  • the maximum sampling rate is typically in the range of about 1 kHz. This results in a maximum depth for the depth scan of about 0.1 mm / s at a center wavelength of the light coupled to the interferometer of, for example, 1300 nm, when four points per period of an interference structure are recorded.
  • the light reflected from the sample 1 and incident on the detector 30 is superimposed with the modulated sensitivity of the detector 30, so that the detector 30 in the detection of the incident on the detector 30 interference pattern instead of a high-frequency interference signal having a plurality of periods a low-frequency Beating signal generated, which has significantly fewer periods than the high-frequency interference signal. Sampling this beat therefore requires significantly fewer sample times per unit time without losing relevant information than sampling the high frequency interference signal without modulating the sensitivity of the detector 30. For a given maximum sample rate of the detector 30, this results in that the maximum speed for a depth scan of the system can be increased many times over.
  • the sensitivity of the detector 30 can be e.g. modulate directly or with a controllable electronic shutter arranged in front of the detector 30.
  • the scan speed is increased by a factor of 4 or even 8.
  • the speed of movement of the reference mirror 16 is preferably in a fixed relationship to the frequency of the modulation of the sensitivity of the detector 30 and is particularly chosen so that in a period of the resulting beat signal an integer number of sampling times, preferably four sampling times match.
  • the beat signals sampled in this way must still be processed before visualization, since the interference information is still contained in these signals.
  • the essential information to be visualized is the amplitude and depth position of the respective interference, but not the interference structure itself.
  • the beat signal must be demodulated by e.g. is determined by Fourier or Hilpert transform the so-called envelope of the beat signal.
  • phase of the beat signal is generally unknown and can also differ for different beat signals from different depths
  • a digital demodulation algorithm is used, which is independent of the phase.
  • so-called 90 ° phase shift algorithms are used for the sampling of the interference signal with four sampling times per period. As a result, a fast demodulation of the beat signal is achieved.
  • a period of modulation of the sensitivity of the detector 30 comprises two sub-periods, during a first sub-period the detector is sensitive and during a second sub-period the detector is insensitive to the light to be detected.
  • the first and second sub-periods are the same length.
  • the first sub-period can be chosen so that its duration is longer than the duration of the second sub-period. In this way, a high signal-to-noise ratio and thus a high image quality is ensured even at low light intensities in addition to a high depth scan speed.
  • the intensity of the light coupled into the interferometer 10 can also be time-modulated, with the explanations regarding the above-described modulation of the detector sensitivity correspondingly applying to the preferred embodiments and the advantageous effects.
  • the radiation source 21 preferably comprises a helical wire, which is surrounded by a transparent envelope, preferably made of glass.
  • the radiation source 21 is formed as a halogen incandescent lamp, in particular tungsten-halogen incandescent lamp, wherein a tungsten wire is used as the wire and the interior of the enclosure is filled with gas containing a halogen, for example iodine or bromine.
  • a tungsten wire is used as the wire and the interior of the enclosure is filled with gas containing a halogen, for example iodine or bromine.
  • the helical wire is made to glow, thereby emitting spatially incoherent light.
  • Spatially incoherent light in the sense of the present invention is to be understood as meaning light whose spatial coherence length is smaller than 15 ⁇ m and in particular only a few micrometers, ie. between about 1 micron and 5 microns, is.
  • the spatially incoherent light generated by the radiation source 21 passes through the optical filter 22, which is designed as a bandpass filter and is substantially transparent only to light within a predefinable spectral bandwidth.
  • the optical filter 22 has a bell-shaped or Gaussian-shaped spectral filter characteristic, wherein only those spectral light components of the light generated by the radiation source 21 can pass through the optical filter 22, which lie within the predetermined bandwidth around a mean wavelength of the bell-shaped or Gaussian spectral filter characteristic ,
  • a Gaussian spectral filter characteristic in the context of the invention means that the transmission of the optical filter 22 for light having specific wavelengths ⁇ is proportional to exp [- [( ⁇ - ⁇ 0 ) / 2 ⁇ ⁇ ] 2 ], where ⁇ 0 the Indicates wavelength at which the optical filter 22 has its maximum transmittance, and ⁇ denotes the standard deviation related to the half width FWHM of the Gaussian transmittance curve, as follows: FWHM ⁇ 2.35 ⁇ ⁇ .
  • a bell-shaped spectral filter characteristic is understood to be a spectral curve of the transmittance of the optical filter 22, which can be approximated by a Gaussian curve and / or differs only so far from a Gaussian curve that its Fourier transform has a substantially Gaussian profile with no secondary maxima or only a few, very low secondary maxima whose height is at most 5% of the maximum of the Fourier transform.
  • thermal radiation sources e.g. Incandescent or halogen lamps are used to generate incoherent light, which are significantly more powerful and less expensive than the commonly used superluminescent diodes (SLDs).
  • SLDs superluminescent diodes
  • the light generated by the radiation source 21 is converted into temporally partially coherent light having a temporal coherence length of preferably more than about 6 ⁇ m.
  • This is of the type described in the described OCT device a so-called time-domain OCT, in which the length of a reference arm 13 in the interferometer 10 changes and continuously, by means of a preferably two-dimensional detector 30, the intensity of the interference is detected, particularly advantageous because the filtering of the light by means of The optical filter 22 realized bandpass achieved on the one hand, a high lateral resolution of the captured image of the sample 1 and the gaussian or bell-shaped spectral filter characteristic of the optical filter 22 on the other hand, an occurrence of disturbing secondary maxima in the Fourier transformation of the interference detected with the detector interference pattern, which Occurrence of further ghosting would be avoided.
  • the described OCT device easily obtains OCT images with high resolution and image quality.
  • the optical filter 22 is arranged between the radiation source 21 and the input-side optics formed by the two lenses 23 and 24. In principle, however, it is also possible to provide the optical filter 22 between the two lenses 23 and 24 or between the lens 24 and the input region 25 of the light guide 26. In principle, an arrangement of the optical filter 22 is particularly advantageous if the light beams impinging on the optical filter 22 have only a small divergence or, in particular, parallel to one another, since on the one hand reflection losses at the interfaces of the optical filter 22 are reduced and on the other hand a beam offset due to refraction of light is reduced. In the example shown, therefore, an arrangement of the optical filter 22 between the two lenses 23 and 24 of the optics is particularly preferred.
  • optical filter 22 directly to the envelope of the radiation source 21. This has the advantage that it is possible to dispense with an additional filter component.
  • optical filter 22 between the output region 27 of the light guide 26 and the illumination arm 12, for example in front of or between the lenses of the optical system 28 located between the output region 27 of the light guide 26 and the input of the illumination arm 12.
  • the optical filter 22 comprises an absorption filter, in particular a so-called mass glass, and an interference filter, wherein several, preferably between about 30 and 70, thin layers with different refractive indices, for example by vapor deposition, are applied to the mass glass, whereby an interference filter is obtained.
  • the optical filter 22 is preferably realized by applying such interference layers to the envelope.
  • the described OCT device can be operated in three different operating modes.
  • the modes of operation are two real-time modes in which OCT images of sample 1 are generated at a high rate of at least one image per second, preferably about 5 to 10 images per second, and a static mode of operation.
  • the real-time mode 1 two-dimensional depth sections of the sample 1 are generated in real time (so-called slices).
  • a CMOS camera is used as the detector 30, which permits the setting of a so-called window of interest (WOI), in which only one Partial surface of the detector 30 is sensitive to light and this converts into corresponding detector signals.
  • WI window of interest
  • the reduction of the sensitive camera area is associated with a significant increase in the camera speed, so that with this setting more camera images can be generated per second than in full-screen mode.
  • a WOI is preferably selected which corresponds in one direction to the total camera length or width (eg 640 pixels) and in the other direction has the minimum possible number of pixels (given by the type of the respective camera) Pixel). This increases the speed of the camera so that OCT images can be captured in real time.
  • FIG. 2 shows as an example a detector 30 with a detector surface A1 which comprises a first number N1 of detector elements 31 arranged in a plane and has a length c1 and a width b1.
  • a detector surface A1 which comprises a first number N1 of detector elements 31 arranged in a plane and has a length c1 and a width b1.
  • the second number N2 of the detector elements 31 of the partial area A2 is smaller than the first number N1 of the detector elements 31 of the entire detector area A1.
  • the lengths c1 and c2 of the detector surface A1 or partial surface A2 are the same, while the widths b1 and b2 of the detector surface A1 or partial surface A2 are different.
  • the partial area A2 is only four pixels wide, whereas the detector area A1 is 512 pixels wide.
  • the sensitive surface of the detector surface A1 is thus reduced by a factor of 128, which considerably shortens the time required for the detection of interference patterns and their conversion into corresponding detector signals.
  • the left part of the FIG. 3 a model of the human skin in which, for example, a plane of a two-dimensional depth slice or slices recorded in operating mode 1, preferably in real time, is drawn.
  • the real-time mode 2 shown - two-dimensional tomograms F generated from a certain depth T of the considered space element R of the sample 1, wherein the depth T is arbitrary.
  • the entire detector surface A1 of the detector 30 is used for detecting the light reflected from the sample 1 and its conversion into corresponding detector signals, but only a maximum of five camera images are used for calculating a tomogram F.
  • the second mode of operation is also referred to as the face mode.
  • the left part of the FIG. 4 a model of the human skin in which, for example, a plane of a two-dimensional tomogram or en-face image recorded in operating mode 2, preferably in real time, is plotted.
  • a complete three-dimensional data set is acquired by means of the macroscopic movement of the reference mirror 16 in combination with focus tracking.
  • the optical path length or amplitude of the macroscopic movement of the reference mirror 16 is at least about 0.1 mm, preferably at least about 1 mm.
  • a macroscopic movement of the reference mirror 16 in the order of 0.1 mm up to several millimeters In contrast to the usual microscopic amplitude of the reference mirror movement in the order of fractions of the average wavelength of the injected light, ie of up to typically 1 .mu.m, takes place in the described OCT device, a macroscopic movement of the reference mirror 16 in the order of 0.1 mm up to several millimeters.
  • the light reflected from the sample 1 is forwarded via the interferometer 10 to the two-dimensional detector 30 and detected successively at a plurality of points in time for a particular period corresponding to the integration time of the detector 30 and into corresponding detector signals transformed.
  • the so-called coherence condition In order for interference to occur between the light reflected by the reference mirror 16 and that reflected by the sample 1, the so-called coherence condition must be satisfied, which i.a. indicates that the respective reflected light waves must have a constant phase relationship with each other in order to be able to interfere with each other. Due to the use of light having a very short coherence length of typically 10 ⁇ m or less, the condition of a constant phase relationship is met only at certain depths or depth ranges of the sample 1, which are also referred to as a coherence gate.
  • Each position of the reference mirror 16 during the macroscopic movement corresponds to a certain depth within the sample 1 or a depth range around this particular depth, for which the coherence condition is satisfied, so that an interference between the reference mirror 16 and the of the Sample 1 reflected light may occur.
  • both half periods of the periodic movement of the reference mirror 16 can each be used to receive detector signals.
  • FIG. 5 illustrates, in which - representative of a plurality of two-dimensional sections - a first, second and third two-dimensional section F1, F2 and F3 is represented by a spatial element R of the sample 1.
  • a two-dimensional section “wanders" in synchronism with the macroscopic movement of the reference mirror 16 in the direction a through the considered spatial element R of the sample 1, without having to move it itself.
  • Each section F1, F2 or F3 lies at a depth T1, T2 or T3 of the sample 1, in each of which the coherence condition is fulfilled, so that an interference between the light reflected by the reference mirror 16 and that of the sample 1 can occur.
  • the macroscopic movement of the reference mirror 16 in combination with the successive two-dimensional detection of the light reflected by the sample 1 thus has the effect of a three-dimensional depth scan.
  • FIG. 5 a model of the human skin, in which, by way of example, a spatial element is shown, from which a three-dimensional tomogram is recorded in operating mode 3.
  • the OCT device described above is designed so that during a full stroke, i. the path length or the double amplitude, the movement of the reference mirror 16 is always an interference signal with sufficiently high intensity and high sharpness is obtained. Moreover, the focus tracking described in more detail below ensures that the interference signal and the sharpness of the detected interference pattern are maximal for all depths in the sample 1.
  • the focus, ie the focal point, of the imaging optics of the interferometer 10 located in the sample arm 14 is set in such a way that the position of the focus in the sample 1 and the position of that plane in the sample 1, in which, in the case of reflection of light, the coherence condition is satisfied and interference occurs, at all times during the acquisition of a tomogram of the spatial element R of the sample 1 are substantially identical. This will be explained below with reference to Figures 6a and 6b illustrated.
  • FIG. 6a shows the case in which the focus f of the sample objective 14a in the sample arm 14, which is shown here only in simplified form, lies in a depth of the sample 1 which does not coincide with the position of the coherence gate K.
  • the section through the sample 1 detected within the coherence gate K at the depth Ti is not exactly focused on the detector 30 (see FIG Fig. 1 ) so that information losses would be tolerated in the detection of the interference.
  • FIG. 6b On the other hand, the case where the focus f of the sample objective 14a has been set to be within the coherence gate K at the depth Ti is shown.
  • This tracking of the focus f of the sample objective 14a corresponding to the respective depth Ti of the coherence gate K is called Focus tracking called.
  • the interferometer 10 is focused on the respective position of the coherence gate K at different depths Ti of the sample 1, so that images of high definition are obtained from each depth of the sample 1.
  • the maximum optical scanning depth Tm indicates to what depth below the surface of the sample 1 the coherence condition for constructive interference is satisfied and corresponding interference patterns are obtained.
  • Sample objective 14a shown in a simplified manner preferably comprises a plurality of lenses which can be moved individually and / or in groups in the direction of sample 1 or away from it.
  • a piezoelectric actuator in particular an ultrasonic piezoelectric motor, is provided, which is coupled to the sample objective 14a or the lenses and moves this or these along one or more guides, in particular guide rods or guide grooves.
  • the movement of the sample objective 14a or of the lenses preferably takes place synchronously with the macroscopic movement of the reference mirror 16 in the interferometer 10 (see FIG Fig. 1 ).
  • the focus f of the sample objective 14a follows the coherence gate G, while the latter passes successively different depths T1, T2 and T3 of the sample 1, from which with the aid of the detector 30 two-dimensional sections F1, F2 and F3 (see FIG , Fig. 5 ).
  • the synchronization of the macroscopic movement of the reference mirror 16 and the focus tracking on the one hand in combination with a two-dimensional detector 30 on the other hand ensures a particularly simple and fast recording a variety of sharp two-dimensional image sections in different depths of the sample 1 and thus the detection of a full three-dimensional image data set with high image quality ,
  • the interference signals detected by the detector 30 are maximal for each depth in the sample 1, resulting in a very high signal-to-noise ratio.
  • this ensures that the lateral resolution is optimal for all depths in the sample 1, since the focus f of the image always lies in the coherence gate K. It preserves true-to-detail, high-contrast OCT images.
  • the speed v2 of the movement of the lens or lenses of the sample objective 14a in the direction of the sample 1 is smaller than the speed v1 of the movement of the reference mirror 16.
  • a ratio v1 / v2 of the speeds of the reference mirror 16 and the lenses is selected is approximately equal to 2 * n-1 or up to about ⁇ 20%, preferably up to about ⁇ 10%, by this value.
  • the ratio v1 / v2 of the speeds of the reference mirror 12 and the lenses 42 ensures that the coherence gate K and the focus f are superimposed during the depth scan over the entire depth range considered.
  • the OCT images obtained with the above-described OCT device or method can be used to further improve the recognition of diagnostic information, for example in the field of dermatology for even better detection of cavities or thickening in the skin with a size of more than about 10 microns, be subjected to an interpolation.
  • One of the OCT images obtained in the above-described OCT device or method, particularly real-time images, is a so-called trilinear interpolation, in which the output image values of at least two two-dimensional output images appear in mutually parallel planes of the Object were recorded, be interpolated in three-dimensional space, so that a two-dimensional final image is obtained. This will be explained in detail below.
  • Trilinear interpolation is a method for multivariant interpolation in a three-dimensional regular lattice, ie a lattice with the same lattice constant in all three spatial directions. This is based on a in FIG. 7 illustrated grid illustrated. From an interpolation of the output image values located at the eight corners C000 to C111 of a cube, a respective interpolation value located at the center C of the cube is derived.
  • the respective output image values are from source images taken in different planes of the object.
  • the output image values are light intensity values at different locations in the corresponding two-dimensional output images.
  • the output image values, ie the light intensity values, with the coordinates C000, C001, C011 and C010 are from a first real-time image recorded in operating mode 1 along a first depth section S (see Fig. 3 ) and the output image values, ie the light intensity values, with the coordinates C100, C101, C111 and C110 from a second real-time image recorded in operating mode 1 along one of the second depth intersections S spaced apart from the lattice constant (see FIG Fig. 3 ).
  • the output image values with the coordinates C000, C010, C110 and C100 originate from a first real-time image recorded in operating mode 2 in the form of a first two-dimensional tomogram F (see FIG Fig. 4 ) and the output image values with the coordinates C001, C011, C111 and C101 from a second real-time image recorded in operating mode 2 in the form of a second two-dimensional tomogram F spaced apart from the lattice constant thereof (see FIG Fig. 4 ).
  • an identical resolution is selected in all three spatial dimensions.
  • trilinear interpolation is the case in the static operating mode both in the case of the two-dimensional real-time images acquired in operating modes 1 and 2 (slice or en-face) and for post-processing 3 obtained three-dimensional tomograms possible.
  • the axial (ie longitudinal) resolution is essentially determined by the spectral bandwidth of the light source 20 and the refractive index of the object 1 to be examined, while the lateral (ie transversal) resolution is essentially determined by the optical imaging and the Size of the detector elements 31 of the detector 30 (see Figures 1 and 2) is determined.
  • the OCT device described above is tuned so that lateral and axial resolution are nearly equal and very high.
  • the resolution in all three dimensions is preferably about 3 ⁇ m ⁇ 3 ⁇ m ⁇ 3 ⁇ m.
  • the depth of field of the imaging optics, in particular the sample objective 14, of the interferometer 10 is greater than the "grid spacing" of the output image values, ie, the spatial distance of the output image values in the three dimensions. This ensures in any case that the output image values are always captured with high accuracy.
  • the fact is taken into account that the sampling of the interference signal must be high enough not to violate the so-called sampling theorem. This will be explained in more detail below.
  • FIG. 8 is a diagram illustrating sampling of an interference pattern 40 in the direction of depth T of an object as compared with the physical resolution 41 in the direction of depth T.
  • An interference period is in this case half the (average) wavelength of the light coupled into the interferometer (at a mean wavelength of approximately 1.3 ⁇ m, this corresponds to approximately 0.65 ⁇ m). It follows that the distance 43 of two sampling points 42 is approximately 0.163 ⁇ m. However, the physical resolution 41 in air is about 4 ⁇ m. This means that approximately 24 consecutive lines in the depth direction T contain approximately the same physical information and can therefore be combined into one line without significant loss of information.
  • the output image value corresponds, for example, to an average value or the median from the original output image values.
  • FIG. 9 Fig. 12 illustrates the summarization of original output image values scanned in the direction of the depth T of the object in several successive lines 44 into a line having only one output image value and one line height, ie a longitudinal extent 45 in the depth direction T corresponding to the lateral extent 46 of a pixel (pixels ) corresponds to the line perpendicular to the depth direction T.
  • detector 30 In operating mode 1, in which slices are recorded in real time, in trilinear interpolation two adjacent rows of the detector 30 are simultaneously read.
  • detector 30 means this is that the width b2 of the partial area A2 of the detector 30 is selected such that it extends in the direction of the width of the detector 30 only over two detector elements 31.
  • the partial area A2 then comprises only 2 x 640 detector elements 31, which during a macroscopic movement of the reference mirror 16 (see FIG FIG. 1 ) are successively read and billed in the manner described above to a two-dimensional final image.
  • the two output images S are taken in the form of two depth sections simultaneously and in a very short time, it is ensured that any relative movement between the sensor head and the object, in particular the human or animal skin, during the recording of the two two-dimensional output images S does not matter.
  • the mid-position reference mirror 16 In operation mode 2, in which en-face images are taken in real time, the mid-position reference mirror 16 (see FIG. 1 ) Only a microscopic, preferably oscillating, movement of about +/- 5 microns to +/- 40 microns.
  • the position or optical imaging property of the sample objective 14 is preferably set so that it has a focal point in a central depth position predetermined by macroscopic displacement of the reference mirror 16.
  • two output images in the form of two en-face images are acquired at two different positions of the reference mirror 16 and formed into a two-dimensional final image of an en-face image.
  • FIG. 11 illustrated diagram showing the course of the position P of the reference mirror 16 over time t.
  • a two-dimensional output image F in the form of a tomogram is obtained from a certain depth in the object by measuring at five positions P of the reference mirror 16 symmetrically about a mean position P 0 .
  • FIG. 11 illustrates the application of trilinear interpolation.
  • Two two-dimensional output images F are obtained by measuring P of the reference mirror 16 at every five positions.
  • the five positions P are in each case symmetrical about the positions P 1 and P 2 , which are preferably themselves symmetrical about the middle position P 0 of the reference mirror 16.
  • the distance 47 of the positions P 1 and P 2 of the reference mirror 16 is in this case by the axial and / or lateral pixel size 45 and 46 (see FIG. 9 ) certainly.
  • the corresponding tomograms F in the object are each half a pixel size above or below the central depth position.
  • the two output images F recorded in this way are then subjected to a trilinear interpolation in which the final image F 'is obtained.
  • the depth of field of the optical image in the interferometer 10 is chosen so that it is greater than half the voxel size.
  • the depth of field must therefore be greater than 1.5 ⁇ m.
  • the sensor head When the images are recorded, the sensor head is preferably in direct contact with the surface of the object to be examined, in particular the skin, which greatly reduces the probability of a relative movement. This is particularly advantageous when taking pictures of human or animal skin, since it is generally elastic and, in particular when a gel is applied, adheres to the tip of the sensor head, so that slight lateral movements or slight tilting of the sensor head are usually not possible cause a relative movement between the skin and the sensor head.
  • operating mode 3 in which static three-dimensional tomograms are recorded, a beat is generated between the detector sensitivity modulation on the one hand and the interference signal to be detected on the other hand, as described in more detail above.
  • the spacing of the individual scanning points in the depth direction is greater than in operating mode 1, so that correspondingly fewer scanning points, preferably between 6 and 10, in particular 8, are combined in order to obtain a cube-shaped three-dimensional image element (voxel).
  • FIG. 12 shows an example of an output image (left) compared to a corresponding end image (right), which was obtained by the described interpolation.
  • the final image is less noisy compared to the original image and therefore appears "softer" or "smoother".
  • the relevant diagnostic information can be obtained more quickly and more reliably from the end images obtained by trilinear interpolation. This applies in particular to cavities or inhomogeneities with a size of typically more than 10 ⁇ m.
  • FIG. 13 shows a schematic representation of a system 50 for carrying out the method according to the invention for optical coherence tomography.
  • the system 50 includes a housing 51, input devices in the form of a keyboard 53, a computer mouse 54 and a footswitch device 55, which has left, middle and right foot switches 55l, 55m and 55r, respectively.
  • the housing 51 is designed to be mobile in the example shown by being provided with rollers 56.
  • a measuring head 57 is provided which is connected to the housing 51 via a cable 58 or a cable hose or tube. In its rest position, the measuring head 57 is inserted in a measuring head holder provided on or in the housing 51, from which it can be removed during the recording of OCT images, which is indicated in the figure by the measuring head 57 shown in dashed lines or the cable 58 shown in dashed lines ,
  • the system comprises a display device 52 in the form of a flat screen on which OCT images 60 and 61, which were recorded by placing the measuring head 57 on an object, in particular the skin of a patient, can be displayed.
  • the first OCT image 60 is a depth cross-section taken substantially perpendicular to the surface of the object being examined, taken in the operating mode 1 described above
  • the second OCT image 61 is a two-dimensional one Tomogram, which runs substantially parallel to the surface of the object under examination and was recorded in the operating mode 2 described above.
  • a straight line 62 is shown in the display device 52, which can be moved in the direction of the indicated double arrow up or down, for example by using the input devices 53, 54 and 55, a corresponding position of the line 62 is selected relative to the first OCT image 60.
  • the system 50 is configured such that, in accordance with the selected position of the straight line 62 in the illustrated first OCT image 60, a plane perpendicular to the first OCT image 60 shown is automatically detected in the examined object and a two-dimensional tomogram is taken there, which is then called the second OCT image 61 is displayed.
  • the first OCT image 60 is preferably a so-called slice, while the second OCT image 61 preferably represents a so-called face image which has been recorded in a plane corresponding to the straight line 62 in the first OCT image 60 ,
  • a depth selection indicator 63 in the form of a switch symbol movable along a straight line which indicates the depth selected by selecting the position of the line 62 relative to the illustrated first OCT image 60.
  • the depth can also be specified in the form of numerical values.
  • a selection display 64 is provided which displays one or more properties of the object to be examined. These properties are preferably selected and entered by an operator prior to taking corresponding OCT images. In dermatological applications, this is, for example, a parameter for characterizing the moisture of the skin of the respective patient.
  • a corresponding switch symbol can then be used along a straight line continuously or in a predetermined manner Steps between the positions "dry skin” left and “wet skin” are moved to the right.
  • interferometer 10 including the optics 28 and the detector 30 integrated.
  • the light source 20 including the input-side optics in the form of the two lenses 23 and 24 are preferably integrated in the housing 51 of the system 50.
  • the light guide 26, by which the light source 20 on the one hand and the interferometer 10 on the other hand are coupled together, is guided in this case within the cable 58 from the housing 51 to the measuring head 57.
  • electrical lines are also performed, which serve on the one hand for power supply and control of the measuring head 57 and on the other hand, the detector signals generated during the detection of OCT images of the detector 30 from this into the housing 51, where they a processing device (not shown ).
  • FIG. 13 only highly schematically shown measuring head 57 is in FIG. 14 shown in detail.
  • a handle 57b is provided, through which the measuring head 57 is removed by an operator from the measuring head holder on or in the housing 51 or inserted again into the measuring head holder and when taking OCT images onto the object placed on and possibly can be performed along this.
  • the measuring head 57 is brought into contact with the object to be examined, in particular the skin of a patient, with a contact surface 57c located at the front end of the measuring head housing 57a.
  • a window 57d is provided, through which light from the sample arm 14 of the interferometer 10 located in the measuring head 57 (see FIG FIG. 1 ) and can thereby irradiate the object to be examined.
  • the light reflected and / or backscattered at different depths of the object re-enters the sample arm 14 of the interferometer 10 through the window 57d and can there, as already detailed above was recorded and evaluated in the form of interference phenomena.
  • a status indicator 57e preferably in the form of a light-emitting display, is further provided, through which e.g. the readiness of the system 50 and / or the measuring head 57 for capturing OCT images is displayed.
  • the cable 58 which may also be designed as a cable channel or hose, is connected to the measuring head 57.
  • the optical coherence tomography system 50 described above three and two-dimensional cross-sectional images of an object, in particular human skin, can be taken, with penetration depths of up to about 1 mm into the human skin and the size of the area of the examined skin area being typical Dimensions of about 1.8 x 1.5 mm. Due to the infrared radiation used in the described system 50 with a preferred mean wavelength of about 1.3 ⁇ m, a radiation exposure of the patient, such as when using X-radiation, can be excluded.
  • the OCT images acquired with the described system 50 also have a high resolution and allow a representation of individual object structures with a size of up to 3 ⁇ m. Last but not least, the OCT images acquired with the system 50 can be used to determine the absolute geometric extent of the different structures, i. their size, to measure.
  • the system 50 comprises a control device for controlling the system 50 according to the invention, in particular the optical coherence tomography device, or for carrying out the processes described above and below.
  • the system further includes a processing means for processing various data including the above-described interpolation of output picture values.
  • the control device and / or the processing device are preferably integrated in the housing 51 of the system 50.
  • FIG. 15 shows the content of the screen 70 of the display device in the management mode, in which the system is automatically after startup.
  • a status indicator 71 in the form of a suitable symbol, for example a green disc, indicates the readiness of the system.
  • the readiness of the system in particular for the acquisition of OCT images, is preferably displayed simultaneously by activation of the status display 57e provided on the measuring head 57.
  • An operator thus has the option of detecting the readiness of the system solely on the basis of the status display 71 on the screen 70 or on the basis of the status display 57e on the measuring head 57.
  • an input field 72 information about the object to be examined, in particular about a patient, can be entered.
  • the system is preferably configured such that a recording of OCT images is only possible if at least one of the information required in the input field 72 is entered, for example at least the last name of a patient.
  • the information entered in the input field 72 in particular first and last name, patient identification number and date of birth, appear then - as in FIG. 16 by way of example - in corresponding fields 72 'in the upper area of the screen display 70.
  • the system automatically starts in operating mode 1 (so-called slice mode).
  • operating mode 1 slice mode
  • an optical gel is applied to the contact surface 57c of the measuring head 57, which on the one hand ensures that sharp refractive index transitions between the skin and the window 57d of the measuring head 57 are bridged (so-called index matching) and on the other hand unevenness on the skin surface are compensated.
  • the amount of optical gel applied to the contact surface 57c is preferably between about 2 ⁇ l and 10 ⁇ l, depending on the application.
  • the contact surface 57c of the measuring head 57 is pressed by an operator against the skin area of the patient to be examined and slightly moved back and forth on the skin area to achieve a favorable distribution of the optical gel.
  • a slice image 73 is detected immediately after establishing a contact with the skin area to be examined and displayed in the region of the center of the screen 70, as in FIG. 17 is illustrated.
  • a display 74 is provided in which the skin type of the respectively examined skin can be set or displayed. This is preferably a parameter characterizing the moisture content of the examined skin area.
  • a corresponding switch symbol can be moved by the operator on a scale between "dry skin” and "moist skin” in stages or also steplessly.
  • this parameter is the ratio of the speeds at which the reference mirror 16 and the lens or lenses of the sample objective 14a (see FIG. 1 such as Figures 6a and 6b ), to ensure optimal focus tracking.
  • Slice image 73 has been selected to have a slightly above the center of the scale position of the switch icon of the display 74, which corresponds to a rather moist skin. As a result, a bright and relatively high-contrast slice image 73 is obtained.
  • FIG. 18 shown slice image 75 at a parameter setting, in which the switch icon of the display 74 is below the center of the scale, which corresponds to a rather dry skin.
  • the contrast of the recorded with this setting slice image 75 against the in FIG. 17 shown Slice image 73 much lower. This can be explained by the fact that when the slice image 75 is recorded at different depths of the skin, the focus of the sample objective 14 has not or not always been located in the region of the respective coherence gate. For further details, reference is made to the above statements in connection with the focus tracking.
  • a slice image 73 obtained with the optimum setting of the parameter corresponding to the skin moisture can be achieved by operating a corresponding switch, preferably by a long press of the middle foot switch 55m (see FIG. 13 ), to switch from slice mode to face mode, in which - as in FIG. 19
  • the slice image 73 in the right-hand area of the screen 70 is shown reduced in size (so-called thumbnail) and at the same time an en-face image 76 recorded in operating mode 2, the so-called face mode, is shown in the middle area of the screen 70 ,
  • the illustrated en-face image 76 is preferably a real-time image that is acquired and updated at a repetition rate of at least one image per second becomes.
  • the reduced representation of the slice image 73 in the right-hand area of the screen 70 is a static image which, for example, is the last one in the slice mode (see FIG FIG. 17 ) in real time and corresponds to the displayed slice image.
  • the depth in the skin in which the displayed en-face image 76 is captured may be selected by the operator via a depth selection switch 77 displayed on the screen 70, by using a corresponding switch icon, for example, by means of the computer mouse 54, the keyboard 53, and / or. or the footswitch device 55 (see FIG. 13 ) is pressed.
  • a depth selection switch 77 displayed on the screen 70
  • the setting or selection of a certain depth by the left foot switch 55l which is designed as a rocker switch and causes by pressing the rocker forward or backward a depth change towards a larger or smaller depth.
  • the system is configured so that a selection of the depth at which an en-face image is to be acquired can be accurate to within a micrometer.
  • the size of the steps in which a depth navigation can be made is specified.
  • the choice of the respective depth for en-face images should be made in increments of 5 ⁇ m. In this way, the depth navigation can be individually adapted to the respective diagnostic purpose.
  • FIG. 20 shows an en-face image 80 which has been taken at a depth which lies between the window 57d on the measuring head 57, which can be seen in the form of a horizontal line 79 in the corresponding slice image 81, and the skin surface Reason only shows a cross section through the gel layer located between the window 57d and the skin.
  • the depth set in this example is displayed by means of a horizontal line 78 drawn in a reduced size in the slice image 81.
  • the selected or set depth of the respective position of the depth selection switch 77 and / or a corresponding numerical value display can be seen.
  • FIG. 12 shows an en-face image 82 taken in a plane lying in the uppermost area of the skin surface, as can be seen in the location of the depth-display line 78 relative to the reduced-size slice image 81.
  • the area between the straight line 78, on the one hand, and the horizontal line 79, which results from light reflections on the window 57d of the measuring head 57, on the other hand corresponds to the gel layer located between the window 57d and the skin.
  • the straight line 78 can be moved relative to the reduced-size slice image 81 (see double arrow), thereby forming planes lying at different depths in the skin can be selected in which appropriate En-face images are recorded and displayed in the screen display.
  • the principle of depth navigation is in the lower right part of the FIG. 21 based on a drawn in a skin model level that is substantially parallel to the skin surface and can be moved in the double arrow direction to different depths, further illustrated.
  • FIG. 22 shows by way of example another en-face image 83, which was taken in a further depth of the examined skin area. As can be seen from the position of the straight line 78 relative to the slice image 81, the plane of the resulting en-face image 83 is now completely within the examined skin area. For the rest, the remarks in connection with the Figures 20 and 21 corresponding.
  • FIG. 23 shows an advantageous use of the depth navigation described above in finding diagnostically relevant information.
  • the choice of the position of the line 78 can be used to select the depth when taking an en-face image 84, for example by assuming a cavity 85 on the basis of the slice image 81 in a corresponding, perpendicular to the Slice image 81 extending level of the en-face image 84 closer.
  • a slice image 85 shown reduced which recorded in slice mode and by pressing a corresponding switch, preferably by briefly pressing the middle foot switch 55m (see FIG. 13 ) was stored in non-volatile memory of the system, eg a hard disk memory.
  • the system is preferably configured such that a further slice image 86 is automatically generated and displayed in the right-hand area of the screen 70 when the slice image last saved, in this case the slice image 85, changes from Slice mode in the En-face mode is already older than a predetermined period of time, for example 10 seconds.
  • a predetermined period of time for example 10 seconds.
  • This configuration of the system ensures that the depth navigation described above is always carried out on the most recent slice image, so that any relative movements between the measuring head on the one hand and the object on the other hand, including movements in the object itself can be considered and thus the reliability of the recording of OCT images, especially en-face images, can not adversely affect.
  • the adjustable period between the recording and storage of a slice image on the one hand and the change from slice mode in the En-face mode on the other hand, when another slice image is recorded, temporarily stored and displayed on the screen 70 was in shown Example set to 10 seconds.
  • this period significantly shorter eg 5 seconds, if this requires the type of the respective examination. This can be the case, for example, if the measuring head can not be held in a fixed position relative to the object for a sufficiently long time due to larger movements of the object, in particular of a patient.
  • FIG. 25 shows a screen display 70 after from the En-face mode, the screen display example in FIG. 24 is shown, was switched back to slice mode.
  • the slice image 85 permanently stored on the basis of a user command is displayed in this case, but not the slice image 86 stored only temporarily for navigation purposes (see FIG FIG. 24 ).
  • the last captured and displayed in en-face mode en-face image 87 is displayed in reduced form.
  • a currently recorded slice image 88 is displayed.
  • the system is configured such that a plane perpendicular to the plane of the illustrated en-face image 87 can also be selected in the reduced-size en-face image 87 with the aid of an additionally represented straight line 89, in which the Slice image 88 is recorded.
  • the principle of lateral navigation is in the lower right part of the FIG. 25 Based on a drawn into a skin model level that is substantially perpendicular to the skin surface and can be moved laterally in the double arrow direction, further illustrated.
  • a slice image 88 ' which was stored in the currently selected slice mode by a corresponding user selection command, is displayed in the right-hand area.
  • a 3D symbol 90 is displayed, which indicates that in the meantime, a recorded in operating mode 3 three-dimensional tomogram has been recorded and stored.
  • the measuring head 57 After completion of the recording of one or more, possibly different, OCT images, the measuring head 57 is inserted again into the measuring head holder located on the housing 51 of the system 50, whereupon the screen display 70 - as shown in FIG. 26 automatically changes to an image viewing mode, in which the operator can select the OCT images 85, 87, 88 'or 90, which are reduced in size, in the right-hand area of the screen 70, the respective selected reduced image being in the region of the center of the screen 70 is shown enlarged.
  • a perspective reproduction of the recorded three-dimensional tomogram can take place in the region of the center of the screen 70.
  • corresponding straight lines 93 and 94 are superimposed into the illustrated slice or face images 91 and 92, respectively.
  • the user can specify the plane of the respectively displayed slice image 91.
  • a selection of a plane of the en-face image from the three-dimensional tomogram to be displayed can be made.
  • the screen 70 shown illustrates how comments can be entered in the image review mode.
  • the operator first selects an OCT image to be commented, in the illustrated example this is the slice image 85, and then a corresponding comment field 97 assigned to this image is opened, in which arbitrary comments can then be entered in the form of free text.
  • a general comment field 96 is opened, into which a comment on the examination performed can be entered, which is assigned to the totality of the OCT images 85, 87, 88 'and 90 recorded in this examination, and when retrieving at least one of these images together with displayed image.
  • the selected in the right area of the screen 70 and reduced shown slice image 85 is shown in enlarged form.
  • a management mode of the system 50 may be selected, in which in the screen view 70, as shown in Figure 28, the performed examination in the form of one line 98 is displayed. By selecting the appropriate line 98, the operator can again switch to the image viewing mode and analyze the recorded OCT images and comment if necessary.
  • an examination report can be drawn up automatically after completion of the examination or after a user command, as shown by way of example in FIG. 29 is shown.
  • the examination report which is preferably generated in HTML format
  • the OCT images 85, 87, 88 'and 90 recorded during the examination and stored on the basis of a user command and the respectively entered comments 96 and 97 in Form summarized.

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EP11002894.1A 2011-04-06 2011-04-06 Verfahren und System zur optischen Kohärenztomographie Not-in-force EP2508842B1 (de)

Priority Applications (5)

Application Number Priority Date Filing Date Title
EP11002894.1A EP2508842B1 (de) 2011-04-06 2011-04-06 Verfahren und System zur optischen Kohärenztomographie
ES11002894.1T ES2497190T3 (es) 2011-04-06 2011-04-06 Sistema y procedimiento para tomografía de coherencia óptica
US14/009,379 US9551565B2 (en) 2011-04-06 2012-03-30 Method and system for optical coherence tomography including obtaining at least two two-dimensional images of an object in three-dimensional space
PCT/EP2012/001436 WO2012136339A1 (de) 2011-04-06 2012-03-30 Verfahren und system zur optischen kohärenztomographie
CN201280017175.2A CN103443578B (zh) 2011-04-06 2012-03-30 用于光学相干断层扫描的方法和系统

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WO2017005838A1 (en) 2015-07-09 2017-01-12 Agfa Healthcare Non-invasive biological tissue examination based on full field high definition optical coherence tomography imaging
US9869541B2 (en) * 2015-07-22 2018-01-16 Medlumics S.L. High-speed optical coherence tomography using multiple interferometers with suppressed multiple scattering cross-talk
US11024004B2 (en) 2018-08-31 2021-06-01 International Business Machines Corporation Reconstructing missing slices in slice-based scans of 3D objects
CN110473285B (zh) * 2019-07-30 2024-03-01 上海联影智能医疗科技有限公司 图像重构方法、装置、计算机设备和存储介质

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ES2497190T3 (es) 2014-09-22
CN103443578B (zh) 2016-07-06
CN103443578A (zh) 2013-12-11

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